I. Field of the Invention
The present invention relates to an imaging apparatus wherein a spin density distribution, relaxation time distribution and chemical shifts of a specified proton (generally, hydrogen nucleus) in biological tissue is measured externally from the object examined (i.e., a patient) in a non-invasive manner by utilizing a nuclear magnetic resonance (NMR) phenomenon so as to obtain information for medical diagnosis. More particularly, the present invention is directed to a magnetic resonance imaging apparatus capable of producing a highly uniform steady magnetic field.
II. Description of the Prior Art
Such an NMR imaging apparatus is described in, e.g., U.S. Pat. No. 4,254,778, issued on Mar. 10, 1981, to Clow et. al.
The known nuclear magnetic resonance techniques (referred to as "NMR" technicues) will be briefly described with reference to FIGS. 1 through 5.
A steady magnetic field H.sub.o is generated by an air coil C1 shown in FIGS. 1A and 1B, and a magnetic gradient field is generated by gradient field generating coils C2, C3 and C4 (FIGS. 2 and 3) assembled together with the air coil Cl. FIG. 4 shows the fields diagrammatically illustrated in the side elevation in relation to a patient P. The steady field H.sub.o generated by the air coil C1 is superimposed in advance on a first gradient field G.sub.z generated by the coils C2. The gradient field G.sub.z can be obtained by flowing reverse currents through a pair of Helmholtz coils C2 shown in FIG. 2. This coil pair is called a "Maxwell pair". The gradient field G.sub.z has the same direction (z-axis) as that of the steady field H.sub.o and has a zero magnetic intensity on a central plane (perpendicular to the z-axis) between the pair of coils C2 so that the absolute values of the intensities of reverse field components linearly increase in opposite directions from the above-described central plane along the z-axis (FIG. 4). The patient P is then placed in the superimposed magnetic field. A selective exciting pulse H.sub.1 having a proper frequency component is applied to the patient at a given time through a pair of saddle-shaped probe head coils C5. The selective exciting pulse H.sub.1 has a center frequency of 4.258 MHz (corresponding to a magnetic field of 1,000 gausses for a hydrogen nucleus) of a carrier wave and is obtained by amplitude-modulating an RF pulse by a SINC function. When the selective exciting pulse H.sub.1 is applied to the patient P, a nuclear magnetic resonance occurs in a plane region (a cross-sectional slice region with respect to the Z axis) wherein a frequency corresponding to a vector sum of the steady field H.sub.o and the gradient field G.sub.z becomes equal to the frequency of the selective exciting pulse H.sub.1. Another magnetic gradient field G.sub.R obtained by a sum of vector components of second magnetic gradient fields G.sub.x and G.sub.y (G.sub.x and G.sub.y are perpendicular to each other and also to G.sub.z) respectively generated by the gradient field generating coils C3 and C4 is applied to the slice S (i.e., chosen slice region) where a nuclear magnetic resonance occurs. In this condition, when a free induction decay signal (referred to as "FID signal") is measured through the probe head coil C5, this signal corresponds to a signal obtained by Fourier-transforming a projection signal indicating a specific nucleus density distribution in the direction of the gradient field GR within the slice S of the patient P. The direction of the gradient field G.sub.R can be varied within the x,y plane by changing the relative ratio of the intensity of the field G.sub.x generated by the coils C3 to that of the field G.sub.y generated by the coils C4. A resultant FID signal is subjected to the inverse Fourier transformation, thereby obtaining projection signals in various directions in the x,y plane. By utilizing these projection signals, an image indicating the density distribution signals, an image indicating the density distribution of the specific nucleus within the slice S of the patient P is obtained.
In general, the slice position of the object under examination such as a patient is very important for an occurrence of the NMR phenomenon. That is, the occurrence probability of the NMR phenomenon owns a significant relation to strengths of the magnetic fields given to the slice region of the object. The specified proton can only resonate with the applied magnetic fields within an extremely narrow field region. It is therefore necessary to generate the highly homogeneous magnetic field in the slice position so as to obtain medical information with better quality, e.g., the high spatial resolution.
Generally, in the diagnostic MNR imaging apparatus, the projection region, i.e., the diagnostic slice region must be less than 100 ppm (10.sup.-4). The higher the uniformity of the magnetic field is, the better the spatial resolution becomes.
In this imaging apparatus, it is very difficult to precisely manufacture the sizes of the main coil (the air coil C1 in FIGS. 1A and 1B) for generating the steady magnetic field and also to precisely define the setting position thereof with respect to the other coils and accessaries. Accordingly, a desirable uniformity of the magnetic field may not be realized by only the main coil, so that the auxiliary coil, so-termed "shim coil" is additionally employed in conjunction with the mail coil. The major function of this slim coil is to correct the precision errors caused by the main coil and also the magnetic disturbances caused by the ferromagnetic materials, e.g., pillar's metals built in the diagnostic examination room, which are located around the main coil.
However, the following difficulty exists in the magnetic field correction by the conventional shim coil.
FIG. 5 schematically shows the conventional shim coil assembly 50, as will be described in detail later, which is wounded on the same bobbin 40 as for the main coil and the gradient field coils. For clearity of illustration, the slim coil assembly 50 is solely shown. The shim coil assembly 50 is constructed by a pair of coil halves 50A and 50B. A radius of the coil halves 50A and 50B is indicated by "a", a diameter thereof being "2a". A distance between the respective coil halves 50A, 50B and a center line CL of the shim coil assembly 50 is denoted by "b", a distance between the opposite sides of the coil halves 50A and 50B being "2b". This shim coil assembly 50 is positioned along the Z axis that is parallel to the longitudinal axis of the object P and also to the direction of the steady magnetic field H.sub.o, and perpendicular to the slice plane of the object P. Coil currents i50A and i50B flows through the respective coil halves 50A and 50B in the directions denoted by arrows. The following Taylor's-formular indicates the nonuniform steady magnetic field Bz in the Z direction, or axis: ##EQU1## where A.sub.i.sup.j, B.sub.i.sup.j (i, j=0, 1, 2 . . . n) is a constant determined by the magnetic field distribution, and X, Y, Z indicates orthogonal coordinates. For the simplicity, the fourth order's term and the succeeding order's terms are omitted.
If a distance measured from an origin in the X - Y coordinate plane is denoted by "R", the following relation can be understood EQU R.sup.2 =X.sup.2 +Y.sup.2 ( 2)
Then the above Taylor's formula (1) can be substituted by the following formula; ##EQU2##
In general, the magnetic field correction by the shim coil, i.e., so-called "shimming", implies the following operation.
In the modified Taylor's formula (3), a plurality of shim coils is provided, each of which generates the specific correction field component for each of the terms (e.g., 3A.sub.2 'X, A.sub.4 .degree.Z(8Z.sup.2 -15R.sup.2)/2) with exception of the constant term, i.e., A.sub.1 .degree., so that the respective terms can be cancelled by the corresponding shim coils. As a result, it can reduce the field strength variations due to the coil's location in the coordinates. Consequently the main coil can generate highly homogeneous steady magnetic field H.sub.o in conjunction with the shimming coils.
On the other hand, if attention is given to the Z direction and the radial direction perpendicular thereto in case of the magnetic field correction by the shim coil, the following problem occurs. When the magnetic field correction for the second order term in the Z direction (axis), the magnetic field in the radial direction is adversely influenced. For instance, if the second order term Z.sup.2 is considered and the following coil data is applied to the shim coil assembly 50 shown in FIG. 5, then the magnetic field distribution is represented in FIG. 6. The respective coil halves 50A and 50B has one turn and the coil current is one ampere. The magnetic density Bz is milli-gauss in unit, and the R and Z axes are plotted in meter. The ratio of the distance "2b" to the coil radius "a" is 0.30 (b/a=0.30).
Under these conditions, the magnetic field correction is performed by the shim coil 50 with respect to the second order term, i.e., 3A.sub.2 .degree.(2Z.sup.2 -R.sup.2)/2 of the formula (3). However, the field distribution in the radial direction is subjected to be distorted. That is to say, as can be seen from the Taylor's formula (3), the magnetic field in the radial (R) direction is substantially uniform, but that in the Z direction is not so homogeneous. If the above-described magnetic field correction by the shim coil 50 is carried out, the magnetic field in the radial direction which has been uniform is necessarily distorted.
Furthermore, due to precision errors in manufacturing the main coil and the shim coil and the magnetic field disturbance by the ferromagnetic materials, the uniformity of the magnetic field in the radial direction cannot be corrected, resulting in the poor quality of the diagnostic images.
An object of the present invention is to provide a magnetic resonance imaging apparatus where highly homogeneous steady magnetic fields can be produced so as to obtain useful medical information by utilizing the nuclear magnetic resonance phenomenon.